The present invention pertains to the field of medical imaging. More particularly, the present invention relates to techniques for correcting for deadtime in a nuclear medicine imaging system.
In nuclear medicine, images of internal structures or functions of a patient""s body are generated by using an imaging system to detect radiation emitted from within the body after the patient has been injected with a radiopharmaceutical substance. The imaging system typically uses one or more scintillator-based detectors to detect the radiation. A computer system generally controls the detectors to acquire data and then processes the acquired data to generate the images. Nuclear medicine imaging techniques include Single-Photon Emission Computed Tomography (SPECT) and Positron Emission Tomography (PET). SPECT imaging is based on the detection of individual gamma rays emitted from the body, while PET imaging is based on the detection of gamma ray pairs that are emitted in coincidence in opposite directions due to electron-positron annihilations. PET imaging is therefore often referred to as xe2x80x9ccoincidencexe2x80x9d imaging. Certain nuclear medicine imaging systems use a small number of (e.g., two) monolithic (continuous) scintillation crystal based detectors, such as dual SPECT/PET systems available from ADAC Laboratories of Milpitas, Calif. Other systems use detectors that consist of a grid of many scintillation crystals, sometimes referred to as xe2x80x9cblock detectorsxe2x80x9d, as with many dedicated PET systems.
One factor that can affect image quality in nuclear medicine imaging systems is deadtime loss. Deadtime can be defined as the inability of a detector to distinguish two distinct scintillation events that occur very close together in time. In other words, deadtime is the time, after detecting an event, during which the detector is xe2x80x9cbusyxe2x80x9d and therefore unable to detect another event. Both the scintillation crystal and the associated electronics may be subject to deadtime. Deadtime loss can be defined as the difference between the true countrate and the observed countrate. FIG. 1 illustrates the effect of deadtime losses in the form of a plot of observed countrate against true countrate. Line 2 represents the ideal yet unrealistic case in which there is no deadtime loss; in that case, the observed countrate OC equals the true countrate C1. In contrast, line 3 represents the response of an imaging system that is subject to deadtime loss; in that case, the observed countrate OC is lower than the true countrate C2. Because deadtime loss is dependent upon the singles rate, deadtime loss increases as the singles rate (true countrate) increases.
One technique for correcting for deadtime loss is to apply a calibration factor to data acquired during an imaging session. For example, deadtime loss can be estimated for a particular imaging system during a pre-clinical calibration session with the use of phantoms. During a clinical imaging session, a correction for the estimated deadtime loss can be applied to the acquired data. One problem with this approach is that deadtime often is not uniform across the imaging surface of the detector, due to variations in the scintillation crystal, parameter of the electronics, and other factors. This problem particularly affects imaging systems that use large, monolithic crystal detectors, which have much larger imaging surfaces than those of block detectors. Also, because deadtime is a function of singles rate, these spatial variations can be exacerbated by the particular energy profile of each patient. Thus, the size, shape, and composition of the patient may contribute to variations in deadtime losses across the imaging surface of the detector. Accordingly, deadtime correction based on pre-clinical calibration and/or the use of phantoms tends to be inaccurate.
An imaging system has a radiation detector which includes a monolithic scintillator and which has deadtime associated with it. A method of correcting for the deadtime includes generating data of an object in response to radiation detected by the detector, correcting the data for spatial variations in deadtime across an imaging area of the monolithic scintillator, and generating an image of the object based on the corrected data.